The present invention relates to class-D output amplifiers for hearing aids.
With the advances in manufacturing technologies and circuit design techniques, hearing aid devices have been considerably reduced in size and their functionality has been increased. In recent years, hearing aid output amplifiers employing class-D architecture have been successfully designed and fabricated. See U.S. Pat. No. 4,689,819 and F. Callias et al., "A Set of Four IC's in CMOS Technology for a Programmable Hearing Aid", IEEE JSSC, p. 301, April 1989.
Class-D amplifiers combine low power consumption with low harmonic distortion levels without having to go to power/distortion tradeoffs of the conventional class-A and class-B amplifiers. Class-A amplifiers offer low distortion with the expense of a high idle current, while class-B amplifiers offer significantly lower idle current operation but suffer from crossover distortion at low signal levels. However, class-D amplifiers are considerably more complicated than their class-A and class-B counterparts.
The traditional method for limiting the output voltage amplitude for class-B hearing aids has been to insert a variable resistor in series with the output transducer. This method is not suitable for class-D output amplifiers.
For a class-B output amplifier there is no current in the load transducer when there is no audio frequency signal to the amplifier. Therefore a resistor in series with the load will not produce any increase in quiescent power consumption.
For the class-D output there are large current flows in the load transducer at all times, even when there is no audio frequency signal. The class-D output circuit relies on the inductive nature of the load transducer to minimize quiescent power consumption. A resistive limiter in series with the output transducer would dissipate steady state power even with no audio frequency signal, thereby defeating the low power consumption advantage of the class-D. A hard limiter scheme for the class-D is needed.
A class-D amplifier uses an oscillator which generates a high frequency triangular waveform. It is also possible to employ an oscillator with a high frequency square wave output. In the latter case, the triangular waveform required in the modulation process is obtained by integrating the square wave. The input to be amplified is added to the triangular wave and this composite signal is compared to a reference voltage level. This process, commonly known as pulse-width-modulation or as pulse-duration-modulation, gives at the output of the comparator a square wave pulse train with a duty cycle continuously changing in response to the input signal. This signal is then provided to a transducer where it is converted back into an amplitude level which is an amplified replica of the input signal.
An important problem of any hearing aid design is stability. The non-zero value of the battery internal impedance causes the voltage at the battery line to be modulated by the signal current flowing through the transducer. This effect can feed back to the front-end circuit blocks where lower signal levels are processed and can cause instabilities. Systems using class-D output amplifiers are no exception in this matter. The class-D output amplifiers of prior art have used compensation schemes to cancel out the adverse effects of the battery internal impedance on the system stability. These compensation schemes are somewhat successful at high frequencies but the proposed methods are not satisfactory at low frequencies. Besides the stability of the amplifier itself, these output amplifiers required external resistor-capacitor (RC) lowpass filters to filter out the unwanted signals from the supply voltage to the external circuit blocks such as preamplifiers, filters and to the microphone. Such RC filters often require a large value capacitor which increases the overall size of the hearing aid device.
The batteries found in today's hearing aid devices are typically single cell, 1.2-1.6 volt batteries. Although great advances have been made in the battery technology, the limited voltage and current capacity of a single cell still poses a major challenge in the hearing aid design field. A successful hearing aid circuit must be able to operate at low voltages and consume a minimal amount of power. Recently, attempts have been made to take advantage of the low current consumption of the CMOS devices operating in switching mode. This limits what can be done with CMOS circuitry in an analog hearing aid system. However, bipolar circuitry which is intrinsically better suited for low voltage analog applications can be employed to design higher performance hearing aid devices while still keeping the power consumption at an acceptable level.